Many medical applications requiring 3D real-time X-ray imaging, like angiographic interventions including vascular stent and stent graft placements, transcatheter embolization and targeted intravascular oncologic procedures, or non-angiographic procedures such as image guided orthopedic, thoracic, abdominal, head and neck, and neuro surgery, biopsies, brachytherapy or external beam radiotherapy, percutaneous drain and stent placement or radiofrequency ablation, could be enhanced by the availability of a faster, more efficient and higher resolution detector. Patients' radiations dose absorption remains obviously a major concern and should be reduced as far as possible.
Also breast imaging requires high resolution at an efficient use of radiation dose for reliable detection of calcifications and delineation of soft tissues still unmet with current imaging technologies. In this field of applications, a trend towards 3D imaging is also observed (see for example I. Sechopoulos in Med Phys. (2013); 40 (1)014302, the entire disclosures of which are hereby incorporated by reference). Tomosynthesis is one of these new imaging technology that fuses CBCT reconstruction with digital image processing to generate images of specified cross sections from a single tomography scan. These devices may as well be much improved with the help of new more efficient detectors.
Flat-panel detectors (FPD) are widely used for medical imaging and have critically increased safety of minimally invasive and endovascular procedures, however, at the price of increased patient and operator irradiation because of limited detection efficiency. There are two fundamentally different designs of FPD. The first (I) is based on indirect conversion, i.e. a two-stage process in which X-rays incident on an absorption layer are transformed to visible photons which are then detected by ordinary photodetectors. The second design (II) rests on the direct conversion of X-rays into electrical signals within a semiconducting absorption layer.
The physics of indirect detectors remains essentially unchanged from that of medical X-ray image intensifiers (XII), which have dominated real time radiographic imaging for over fifty years. The conversion of X-ray photons into visible photons takes place in a scintillation layer such as CsI(TI). Apart from limited spatial and energy resolution, the two-stage conversion process suffers from the drawback of low conversion efficiency and limited spectral resolution. The efficiency of wavelength conversion in CsI(TI), for example, may be around 10% and the subsequent conversion of the visible light into electron-hole (e-h) pairs in the photodetectors typically has an efficiency below 50%. As a result, the number of electron-hole pairs collected by the readout circuit is on the order of 25 per keV of X-ray energy (see U.S. Pat. No. 8,237,126 to H. von Känel, the entire disclosure of which is hereby incorporated by reference).
These drawbacks are essentially absent in direct detection of X-rays by means of semiconductor absorbers in which X-rays are converted into electron-hole (e-h) pairs giving rise to electrical signals. For reasons of manufacturability and costs, current devices are, however, based on polycrystalline or amorphous materials and readout circuits made from thin film transistors. Such FPD can be produced in a monolithic form, wherein the absorber layer is deposited directly on the readout electronics in a low temperature process. The only FPD using a direct conversion mode and are commercially available for medical applications are in fact based on amorphous selenium (a-Se) absorbers. They offer large size and are relatively inexpensive to make (see for example S. Kasap et al. in Sensors 11, 5112 (2011), the entire disclosure of which is hereby incorporated by reference). The detective quantum efficiency achievable with polycrystalline and or amorphous semiconductors is, however, again rather low because of their poor electrical transport properties, making FPD derived from such materials ill-suited for interventional radiology.
By far the best direct X-ray imaging detectors for what concerns spatial, temporal and energy resolution as well as detective quantum efficiency are those based on single crystal absorbers. The list of materials suitable for the fabrication of single crystalline X-ray absorbers is rather limited because they need to have a suitable bandgap and must be of sufficient size and perfection. The list includes mainly silicon (Si), germanium (Ge), gallium arsenide (GaAs), cadmium telluride (CdTe) and cadmium zinc telluride (CdZnTe).
Semiconductor absorbers based on any of these materials offer spectral resolution, since the energy of an incident X-ray photon is proportional to the number of generated e-h pairs and thus measurable by a pulse height analysis in single-photon counting.
In Si, one needs on average 3.6 eV to create a single e-h pair (see for example R. C. Alig et al. in Phys. Rev. B 22, 5565 (1980); and R. C. Alig in Phys. Rev. B 27, 968 (1983), the entire disclosures of which are hereby incorporated by reference). On average this leads to 280 e-h pairs per keV of absorbed X-ray energy, from which it can be seen that the conversion efficiency exceeds that of a scintillator-photodiode combination by more than a factor of ten. On the other hand, its low atomic number Z, makes Si a poor absorber for photon energies above about 20 keV. Moreover, the Compton Effect competing with the photoelectric effect at high photon energies (>57 keV) tends to offset the energy resolution in photon-counting. According to recent studies, this problem can, however, be solved (see for example U.S. Pat. No. 8,378,310 to Bornefalk; and H. Bornefalk et al. in Phys. Med. Biol. 55, 1999 (2010), the entire disclosures of which are hereby incorporated by reference). Similarly, the problems of cross-talk between neighboring pixels and signal pileup for high count rates appear to be solvable (see for example U.S. Pat. No. 8,378,310 to Bornefalk; and Bornefalk et al. in Nucl. Instr. Meth. Phys. Res. A 621, 371 (2010), the entire disclosures of which are hereby incorporated by reference).
In the meantime, Si-strip detectors with energy discriminating capabilities have reached the commercial stage for breast imaging (see H.-M. Cho, Med Phys. 41 091903 (2014), the entire disclosure of which is hereby incorporated by reference). Segmented Si-strip detectors in edge-on geometry to cover most of the photon energy range used in computed tomography (CT) are under development (see for example X. Liu et al. in IEEE Trans. Nucl. Sci. 61, 1099 (2014), the entire disclosure of which is hereby incorporated by reference).
In the usual detector geometry, such as that of a FPD, the use of Si is limited to low photon energies because of its low Z. The only elemental semiconductor with a higher Z for which large wafers of excellent quality are commercially available is Ge. It has, however, the drawback of a small bandgap Eg of only 0.66 eV, leading to a resistivity of only 50 Ωcm at room temperature, even when it is undoped and highly pure. As a result, Ge detectors need to be cooled, at least to about −50° C. in order to reduce the excessive dark current related to the low resistivity (see for example D. Pennicard et al. in Jinst 6, C11009 (2011), the entire disclosure of which is hereby incorporated by reference). Higher room temperature resistivities are offered by SiGe alloys up to the Ge content (around 80%) at which the nature of the bandgap changes from Si-like to Ge-like (see for example J. Weber et al. in Phys. Rev. B 40, 5683 (1989), the entire disclosure of which is hereby incorporated by reference). Since, however, large SiGe wafers are not commercially available, one has to resort to epitaxial growth on Si substrates in order to realize FPD based on SiGe. The large mismatch of lattice parameters and thermal expansion coefficients of Ge and Si cause, however, high defect densities (such as misfit and threading dislocations and stacking faults) and cracks in an epitaxial SiGe layer of sufficient thickness on the order of 100 μm or larger to serve as an X-ray absorber. In addition, device processing may not be permitted at all due to excessive wafer bowing resulting from the thermal misfit.
The problem of wafer bowing and layer cracking has been solved by a method involving deep Si-substrate patterning at a micron-scale, along with far-from-equilibrium epitaxial growth. This gives rise for example to space-filling, three-dimensional (3D) SiGe-crystals separated by tiny gaps (see for example International Patent Application No. WO 2011/135432 to H. von Känel; and C. V. Falub et al. in Science 335, 1330 (2012), the entire disclosures of which are hereby incorporated by reference). For sufficiently large aspect ratio of the crystals, provided that they exhibit faceted surfaces, the method leads furthermore to the expulsion of all threading dislocations, so that crystal regions at a distance of several microns from the interface are entirely defect-free (see for example C. V. Falub et al. in Sci. Rpts. 3, 2276 (2013), the entire disclosure of which is hereby incorporated by reference). The approach does not, however, eliminate the high density of misfit dislocations present at the SiGe/Si interface. In order to eliminate this deficiency, the Ge content in the 3D SiGe crystals must be slowly increased from zero up to some final value. This method of compositional grading has been used in the past to lower the threading dislocation density in epitaxial SiGe/Si films, but without affecting the misfit dislocation density (see for example E. A. Fitzgerald et al. in Appl. Phys. Lett. 59, 811 (1991), the entire disclosure of which is hereby incorporated by reference). In the tall 3D SiGe crystals suitable for X-ray absorbers, one can expect the misfit stress to be relaxed elastically provided that the grading rate is kept low enough. Because of the lack of misfit dislocations these structures should therefore be entirely defect-free (see for example M. Salvalaglio, J. Appl. Phys. 116, 104306 (2014), and F. Isa et al. in Acta Materialia 114, 97 (2016), the entire disclosures of which are hereby incorporated by reference).
While epitaxial SiGe absorbers can be grown on large wafers at least 200 mm in size, they are necessarily limited in thickness to a range of about 100-200 μm for example for cost reasons. Under these conditions, efficient X-ray absorption close to 100% is found only for tube voltages below about 35 keV. In medical applications such absorbers are therefore best suited for mammography.
Obviously, there is no such thickness limitation for bulk Ge wafers which for 1 mm and 2 mm absorb close to 100% of the radiation emitted by tubes operated at 40 keV and 50 keV, respectively. The X-ray absorption of GaAs is very similar to that of Ge because of similar Z, and large wafers are commercially available as well. The bandgap of GaAs is ˜1.4 eV and it can be made semi-insulating with a resistivity up to 109 Ωcm, such that cooling is not necessary (see for example M. C. Veale et al. in Nucl Instr. Meth. Phys. Res. A 752, 6 (2014), the entire disclosure of which is hereby incorporated by reference). On the other hand the mobility-lifetime product is much inferior to that of elemental semiconductors such as Si and Ge, which negatively affects the charge collection efficiency and the energy resolution of detectors made from this material.
In order to compensate for the decreasing charge collection efficiency at large absorber thickness, a three-dimensional detector structure has therefore been conceived, in which electrodes are drilled into the absorber volume, permitting electron-hole pairs to be collected by lateral transport (see for example E. Gros d'Aillon et al. in Nucl. Instr. Meth. Phys. Res. A 727, 126 (2013), the entire disclosure of which is hereby incorporated by reference).
In principle, CdTe and CdZnTe with a Zn content of about 10% should be the best materials for efficient X-ray absorption, especially for the higher photon energies required for CT (up to about 140 keV), due to their even higher Z. These are II-VI semiconductors with a higher degree of iconicity than III-V semiconductors, such as GaAs mentioned above. They contain multiple deep trap levels which affect especially the hole transport, leading to the phenomenon of polarization in which positive charges pile up in front of the cathode under high X-ray flux levels (see for example D. S. Bale et al. in Phys. Rev. B 77, 035205 (2008), the entire disclosure of which is hereby incorporated by reference). This strongly modifies the electric field distribution inside the absorber and negatively affects the charge collection efficiency and the measured energy spectrum in the photon counting mode of the detector. Other effects degrading detector performance are pulse pileups (more severe compared to GaAs because of lower carrier mobility), charge sharing among neighboring pixels, fluorescence producing lower energy photons, and Compton scattering (see for example K. Taguchi et al. in Medical Physics 40, 100901 (2013), the entire disclosure of which is hereby incorporated by reference). Presently, perhaps the biggest disadvantage is, however, the lack of large high-quality wafers which creates a tiling issue that doesn't allow for imagine without dead spaces or makes these materials ill-suited for the fabrication of FPD and applications where a high-energy resolution is required.
Irrespective of the kind of material used for the X-ray absorber, the latter needs to communicate with the readout electronics by means of which the analog charge pulses generated by the absorbed photons can be amplified, shaped and transformed into digital signals. The fabrication of monolithic structures by direct deposition of the absorber onto the readout wafer, discussed above for polycrystalline and amorphous absorber materials, is not possible for single crystal absorbers because of the high thermal budget required by epitaxial (single crystalline) growth, the large lattice misfits causing high defect densities, and the thermal mismatch responsible for wafer bowing and layer cracking. Epitaxial SiGe absorbers grown in the form of tall crystals on patterned CMOS processed Si substrates may be the only exception offering sufficient material quality, but thermal budget constraints are present here as well and require a special high temperature metallization scheme for the readout circuits, so that state-of-the-art CMOS processing cannot be used (see for example U.S. Pat. No. 8,237,126 to von Känel, the entire disclosure of which is hereby incorporated by reference).
The only monolithic pixel sensors with single crystal absorbers available to date are designed for particle detection in high-energy physics, where the Si absorption layer does not need to be thick (see for example S. Mattiazzo et al. in Nucl. Instr. Meth. Phys. Res. A 718, 288 (2013), the entire disclosure of which is hereby incorporated by reference).
In practice, establishing the electrical connections through which the charge pulses generated in the absorber by incident X-ray photons are transmitted to the readout unit requires some form of bonding process between the two. The common bonding technique for two-dimensional detectors used today is bump bonding, as for example employed by the Medipix collaboration, CERN (http://medipix.web.cern.ch) or by Dectris AG (http://www.dectris.ch, of Baden-Daettwil, Switzerland). This hybrid approach is very flexible, applicable in principle to any semiconductor material suitable for X-ray detection of which sufficiently large single crystals are available (see for example European Patent No. 0571135 to Collins et al., the entire disclosure of which is hereby incorporated by reference). Such bump-bonded detectors find numerous applications in biology (see for example E. Bertolucci in Nucl Sci. Meth. Phys. Res. A 422, 242 (1999), the entire disclosure of which is hereby incorporated by reference); material science (see for example R. Ballabriga et al. in Jinst 8, C02016 (2013), the entire disclosure of which is hereby incorporated by reference), including CT in material science (see for example S. Procz et al. in Jinst 8, C01025 (2013), the entire disclosure of which is hereby incorporated by reference). Very recently, a Medipix3RX spectroscopic pixel detector with a GaAs absorber has been introduced and shown to be suitable for spectroscopic CT acquisitions, and probably soon for small animal imaging (see for example E. Hamann et al. in IEEE Trans. Med. Imaging 34, 707 (2015), the entire disclosure of which is hereby incorporated by reference).
The bump bonding technique is, however, rather expensive and therefore ill-suited for the fabrication of large area detectors (FPD) such as the ones needed for example for Cone Beam Computed Tomography (CBCT). In order for the single readout chips to be buttable, the through-silicon-via (TSV) technology originally developed for the 3-dimensional integration of microelectronics chips must be used, adding to the complexity of bump bonding (see for example Z. Vykydal et al. in Nucl. Instr. Meth. in Phys. Res. A 591, 241 (2008) and D. Henry et al. in IEEE Electronic Components & Technology Conference 2013, pp. 568, the entire disclosures of which are hereby incorporated by reference). One way to overcome these limitations would be to replace the expensive bump bonding technique by direct wafer bonding, wherein the absorber wafer is bonded to the wafer containing the readout circuits. Detectors made by direct bonding of readout and absorber wafers can no longer be distinguished from those having epitaxial readout/absorber interfaces and therefore may be termed “monolithic” with equal justification as detectors featuring absorber layers directly grown onto the readout wafer. The fabrication of detectors by means of direct wafer bonding requires a low-temperature wafer bonding technique for which several approaches have become available only recently.
The first approach is based on hydrophobic bonding of two hydrogen passivated wafers. It has been shown to be applicable to the direct bonding of two Si wafers, each of which has been device processed prior to the bonding. Hydrophobic bonding has the advantage of providing electrically transparent interfaces because the native oxides have been removed by the H-passivating step. The interfacial hydrogen has to be removed after bonding, however, in order to result in high bond strength. This can be achieved by thermal annealing above the hydrogen desorption temperature, which is a critical step because of thermal budget constraints which do not allow temperatures above 450° C. for standard aluminum metallization. The annealing has the further disadvantage of gas evolution at the interface, gas bubbles causing electrically insulating voids. In order to remove such gas bubbles, trenches may be etched prior to the bonding (see for example U.S. Pat. No. 6,787,885 to Esser et al., the entire disclosure of which is hereby incorporated by reference). The main disadvantage of hydrophobic remains, i.e. the required removal of interfacial hydrogen by thermal annealing. In particular, this rules out the bonding of wafers differing in thermal expansion coefficients, such as for example bonding of Si to SiGe, GaAs or CdTe.
What is needed is therefore a low-temperature process resulting in electrically transparent direct wafer bonds of bulk strength without the need of any high temperature annealing step. Bulk bond strength has been achieved in a process of covalent wafer bonding developed for example by EV Group for Si-wafer bonding at temperatures below 100° C. or even as low as room temperature (see for example C. Flötgen et al. in ECS Transactions 64, 103 (2014), the entire disclosure of which is hereby incorporated by reference). Electrical transport experiments have shown, however, that the oxide removal by dry etching prior to covalent bonding may result in interfacial barriers hindering charge carriers to cross the bonding interface. Surface amorphization during oxide removal must therefore be avoided or amorphous layers be passivated for example by a hydrogen passivation step known in the art (see for example A. Loshachenko et al. in Phys. Stat. Sol. C 10, 36 (2013), and T. Jiang et al. in Phys. Stat. Sol. A 209, 990 (2012), the entire disclosures of which are hereby incorporated by reference).
This invention may be applied to the fabrication of large area monolithic pixel sensors even for high-Z absorber materials for which at present no large wafers can be manufactured at a bearable cost.
What is needed is Computed Tomography (CT) equipment which overcomes the excessive radiation exposure to which patients are subjected by present day FPD. Current energy integrating FPD are therefore to be replaced by FPD with energy discriminating capabilities, high sensitivity and superior spatial resolution. The use of the photon counting mode offers substantially lower radiation doses for a wide range of medical imaging procedures.
A CBCT unit may comprise at least one X-ray source and a FPD mounted on a gantry disk, allowing the patient to remain stationary during the examination (see for example R. Baba et al. in Comp. Med. Imaging and Graphics 26, 153 (2002), and R. Gupta et al. in Eur. Radiol. 16, 1191 (2006), the entire disclosures of which are hereby incorporated by reference). Alternatively, C-arm mounted CBCT units are ideally suited for imaging in the interventional suite (see for example S. Hirota et al. in Cardiovasc. Intervent. Radiol. 29, 1034 (2006), and R. C. Orth et al. in J. Vasc. Interv. Radiol. 19, 814 (2008), the entire disclosures of which are hereby incorporated by reference). C-arm mounted CBCT units employing such energy resolving digital FPD will allow real-time 2D tissue-specific imaging and volumetric data acquisition in a single rotation of the source and the FPD. The photon counting mode will permit the reduction of either the amount of contrast agent or the radiation dose while maintaining the contrast-to-noise ratio at the level of current FPD. As demonstrated by K. Taguchi et al. (in: Med Phys. October 2013, the entire disclosure of which is hereby incorporated by reference), the contrast dose may be reduced by e.g. 23% or the radiation dose by 41%. Si-based photon counting detectors can even contribute to an average dose reduction of approximatively 40%, as demonstrated with the Philips MicroDose SI mammography system (Philips Healthcare, White paper, 2012. Comparison of Dose Levels in a National Mammography Screening Program).